This invention relates to X-ray computed tomography (CT) imaging apparatus such as used in medical imaging; and more particularly to systems for calibrating the detector array in such apparatus.
In computed tomography, an X-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a cartesian coordinate system. The X-ray beam passes through an object being imaged, such as a medical patient, and impinges upon an array of radiation detectors. The intensity of the transmitted radiation is dependent upon the attenuation of the X-ray beam by the object and each detector produces a separate electrical signal that is a measurement of the beam attenuation. The attenuation measurements from all of the detectors are acquired separately to produce a transmission profile.
The source and detector array in a conventional CT system are rotated on a gantry within the imaging plane and around the object so that the angle at which the X-ray beam intersects the object constantly changes. A group of X-ray attenuation measurements from the detector array at a given angle is referred to as a "view" and a "scan" of the object represents a set of views made at different angular orientations during one revolution of the X-ray source and detector. The data for a given scan is stored in memory as a two-dimensional array with storage locations along one axis of the array representing the data from each radiation detector and the storage locations along the other axis containing data for the different views.
The scan data is processed to construct an image that corresponds to a two-dimensional slice taken through the object. The prevailing method for reconstructing the image is referred to in the art as the filtered backprojection technique. This process converts the attenuation measurements from a scan into integers called "CT numbers" or "Hounsfield units," which are used to control the brightness of a corresponding picture element on a cathode ray tube display.
In order for an image to be accurately constructed from the scan data, the signal processing circuitry for each detector in the array must be calibrated so that when a uniform level of radiation impinges each detector, the same magnitude signal will be produced from every processing circuit for the detectors.
Heretofore, the common technique for calibrating the detector array and processing circuits was referred to as the "air calibration technique." In performing this technique, the X-ray source was activated to produce a known intensity X-ray beam that was directed toward the detector array without any object being present therebetween. Ideally, this technique should produce the same level output signal from the processing circuit for each radiation detector. Any deviation of the signal of a given detector from a nominal level represented an error for that particular detector element. That erroneous signal level then was used in adjusting the gain of the signal processing circuit for the associated detector to produce the nominal output level during the air calibration.
If the detector arrays were relatively stable, the air calibration process needed to be only performed once a day. However, certain types of detector elements degraded with continuing exposure to X-rays. For these types of detector arrays, the air calibration technique had to be performed prior to imaging each patient. Even when air calibration needed to be performed only once a day, it consumed time that could otherwise be used for patient imaging. Therefore, it is desirable to reduce the frequency at which the detector array and its associated signal processing circuitry needs to be calibrated.